Method of making a poly(l-lactide) stent with tunable degradation rate

ABSTRACT

Methods of treating with a biodegradable polymeric stent made from poly(L-lactide) and a low concentration of L-lactide monomer is disclosed. The concentration of L-lactide is adjusted to provide a degradation behavior that is suitable for different treatment applications including coronary, peripheral, and nasal.

CROSS-REFERENCE

This application is a continuation of application Ser. No. 14/970,890filed Dec. 16, 2015, which is a continuation of application Ser. No.14/530,595, now U.S. Pat. No. 9,248,218, which was filed on Oct. 31,2014, which is a continuation of application Ser. No. 12/751,773, nowU.S. Pat. No. 8,889,823, which was filed on Mar. 31, 2010, which is acontinuation-in-part of application Ser. No. 12/506,973 filed on Jul.21, 2009, all of which are incorporated by reference herein.

BACKGROUND OF THE INVENTION

Field of the Invention

This invention relates to methods of treatment of blood vessels withbioabsorbable polymeric medical devices, in particular, stents.

Description of the State of the Art

This invention relates to radially expandable endoprostheses, that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, formed fromwires, tubes, or sheets of material rolled into a cylindrical shape.This scaffolding gets its name because it physically holds open and, ifdesired, expands the wall of the passageway. Typically, stents arecapable of being compressed or crimped onto a catheter so that they canbe delivered to and deployed at a treatment site.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance. Thetherapeutic substance can also mitigate an adverse biological responseto the presence of the stent. Effective concentrations at the treatedsite require systemic drug administration which often produces adverseor even toxic side effects. Local delivery is a preferred treatmentmethod because it administers smaller total medication levels thansystemic methods, but concentrates the drug at a specific site. Localdelivery thus produces fewer side effects and achieves better results.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesan active or bioactive agent or drug. Polymeric scaffolding may alsoserve as a carrier of an active agent or drug.

The stent must be able to satisfy a number of mechanical requirements.The stent must have sufficient radial strength so that it is capable ofwithstanding the structural loads, namely radial compressive forces,imposed on the stent as it supports the walls of a vessel. Onceexpanded, the stent must adequately provide lumen support during a timerequired for treatment in spite of the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.In addition, the stent must possess sufficient flexibility with acertain resistance to fracture.

Stents made from biostable or non-erodible materials, such as metals,have become the standard of care for percutaneous coronary intervention(PCI) as well as in peripheral applications, such as the superficialfemoral artery (SFA), since such stents have been shown to be capable ofpreventing early and later recoil and restenosis.

In order not to affect healing of a diseased blood vessel, the presenceof the stent is necessary only for a limited period of time. There arecertain disadvantages to the presence of a permanent implant in a vesselsuch as compliance mismatch between the stent and vessel and risk ofembolic events. To alleviate such disadvantages, stent can be made frommaterials that erode or disintegrate through exposure to conditionswithin the body. Thus, erodible portions of the stent can disappear fromthe implant region after the treatment is completed, leaving a healedvessel. Stents fabricated from biodegradable, bioabsorbable, and/orbioerodable materials such as bioabsorbable polymers can be designed tocompletely erode only after the clinical need for them has ended.

Like a durable stent, a biodegradable stent must meet time dependentmechanical requirements. For example, it must provide patency for aminimum time period. However, it is also important for a biodegradablestent to completely degrade from the implant site with in a certainperiod of time. A biodegradable material that can provide the mechanicalrequirements may not possess the requisite or desired degradation time.In addition, the requisite or desired degradation time varies betweentypes of applications, i.e. coronary or peripheral.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method offabricating a stent comprising: making PLLA polymer material with apolymerization reaction of LLA monomers, wherein the PLLA polymermaterial has an LLA monomer content of greater than 4 wt %; processingthe PLLA polymer material to reduce the LLA monomer content to no morethan 2 wt % and; making a polymeric stent body from the processed PLLA.

Additional embodiments of the present invention include a method offabricating a stent comprising:_obtaining a PLLA polymer material havingan LLA monomer content of greater than 4 wt %; processing the PLLApolymer to reduce the monomer content to no more than 2 wt %; and makinga polymeric stent body from the processed PLLA.

Other embodiments of the present invention include a method offabricating a stent comprising: making a PLLA polymer material with apolymerization reaction of LLA monomer, wherein the PLLA polymermaterial has a LLA monomer content of less than 0.5 wt %; increasing thereaction temperature of the PLLA polymer material to causedepolymerization which increases the monomer content to a target LLAmonomer content of between 0.5 to 2 wt % of the PLLA; and making a stentbody from the PLLA polymer material with the increased monomer content.

Further embodiments of the present invention include a method of makinga stent comprising: selecting a reaction temperature between 120-180° C.of a polymerization reaction of LLA monomer to PLLA; selecting a targetLLA monomer content for PLLA polymer material from the polymerizationreaction between 0.5 to 2 wt %; performing the polymerization reactionfor different reaction times to determine a target reaction time thatprovides a PLLA polymer material with the target LLA monomer content;making the PLLA polymer material with the target LLA monomer contentfrom the polymerization reaction at the target reaction temperature forthe target reaction time; and making a stent from the PLLA polymermaterial with the target LLA monomer.

Additional embodiments of the present invention include a method ofmaking a stent comprising: selecting a reaction time between 1-100 h ofa polymerization reaction of LLA monomer to PLLA; selecting a targetmonomer content of PLLA polymer material from the polymerizationreaction between 0.5 to 2 wt %; performing the polymerization reactionfor different reaction temperatures between 120-180° C. to determine atarget reaction temperature that provides the target LLA monomercontent; making the PLLA polymer material with the target LLA monomercontent from the polymerization reaction at the target reactiontemperature for the target reaction time; and making a stent from thePLLA polymer material with the target monomer content.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts an exemplary stent.

FIG. 2 depicts a plot of the in vitro degradation behavior for a PLLAstent with different concentrations of L-lactide monomer.

FIG. 3 depicts a plot of a degradation rate constants versus L-lactidemonomer concentration based on the data of FIG. 2.

FIG. 4 depicts the dependence of the monomer conversion on the reactiontime for bulk polymerization of L-lactide at different temperatures.

DETAILED DESCRIPTION OF THE INVENTION

Coronary arteries refer generally to arteries that branch off the aortato supply the heart muscle with oxygenated blood. Peripheral arteriesrefer generally to blood vessels outside the heart and brain.

In both coronary artery disease and peripheral artery disease, thearteries become hardened and narrowed or stenotic and restrict bloodflow. In the case of the coronary arteries, blood flow is restricted tothe heart, while in the peripheral arteries blood flow is restrictedleading to the kidneys, stomach, arms, legs, and feet. The narrowing iscaused by the buildup of cholesterol and other material, called plaque,on the inner walls of the vessel. Such narrowed or stenotic portions areoften referred to as lesions. Artery disease also includes thereoccurrence of stenosis or restenosis that occurs after an angioplastytreatment. Although there are probably several mechanisms that lead torestenosis of arteries, an important one is the inflammatory response,which induces tissue proliferation around an angioplasty site. Theinflammatory response can be caused by the balloon expansion used toopen the vessel, or if a stent is placed, by the foreign material of thestent itself.

Embodiments of the present invention are applicable to treatment ofcoronary and peripheral disease in coronary arteries and variousperipheral vessels including the superficial femoral artery, the iliacartery, and carotid artery. The embodiments are further applicable tovarious stent types, such as self-expandable and balloon expandablestents. The embodiments are further applicable to various stent designsincluding scaffolding structures formed from tubes, wire structures, andwoven mesh structures.

In embodiments of the present invention, a stent includes a plurality ofcylindrical rings connected or coupled with linking elements. Whendeployed in a section of a vessel, the cylindrical rings are loadbearing and support the vessel wall at an expanded diameter or adiameter range due to cyclical forces in the vessel. Load bearing refersto the supporting of the load imposed by radial inwardly directedforces. Structural elements, such as the linking elements or struts, arenon-load bearing, serving to maintain connectivity between the rings.For example, a stent may include a scaffolding composed of a pattern ornetwork of interconnecting structural elements or struts.

FIG. 1 depicts a view of an exemplary stent 100. In some embodiments, astent may include a body, backbone, or scaffolding having a pattern ornetwork of interconnecting structural elements 105. Stent 100 may beformed from a tube (not shown). FIG. 1 illustrates features that aretypical to many stent patterns including cylindrical rings 107 connectedby linking elements 110. As mentioned above, the cylindrical rings areload bearing in that they provide radially directed force to support thewalls of a vessel. The linking elements generally function to hold thecylindrical rings together.

The structural pattern in FIG. 1 is merely exemplary and serves toillustrate the basic structure and features of a stent pattern. A stentsuch as stent 100 may be fabricated from a polymeric tube or a sheet byrolling and bonding the sheet to form the tube. A tube or sheet can beformed by extrusion or injection molding. A stent pattern, such as theone pictured in FIG. 1, can be formed on a tube or sheet with atechnique such as laser cutting or chemical etching. The stent can thenbe crimped on to a balloon or catheter for delivery into a bodily lumen.

The prevailing mechanism of degradation of biodegradable polymers ischemical hydrolysis of the hydrolytically unstable backbone. In a bulkeroding polymer, polymer is chemically degraded and material is lostfrom the entire polymer volume. As the polymer degrades, the molecularweight decreases. The reduction in molecular weight is followed by areduction in mechanical properties, and then erosion or mass loss. Thedecrease in mechanical properties eventually results in loss ofmechanical integrity demonstrated by fragmentation of the device.Enzymatic attack and metabolization of the fragments occurs, resultingin a rapid loss of polymer mass.

The treatment of artery disease with a stent of the present inventionhas time dependent properties once it is implanted which enable thetreatment and healing of a diseased section of the vessel. Inparticular, the molecular weight, the mechanical properties, themechanical integrity, and mass change with time. After deployment at adiseased section artery, the stent supports the section at an increaseddiameter for a period of time. Due to a decrease in molecular weight,the radial strength degrades to the point that the stent can no longersupport the walls of the section of the vessel. “Radial strength” of astent is defined as the pressure at which a stent experiencesirrecoverable deformation. The loss of radial strength is followed by agradual decline of mechanical integrity.

Mechanical integrity refers to the size, shape, and connectivity of thestructural elements of the stent. For example, the shape refers to thegenerally tubular shape of the stent formed by the cylindrically-shaperings connected by the linking elements of the pattern. Mechanicalintegrity starts to be lost when fractures appear or propagate instructural elements of the stent due to chemical degradation (molecularweight decline). Further loss of mechanical integrity occurs when thereis breaking or loss of connectivity in structural elements.

The initial clinical need for any stent is to provide mechanical supportto maintain patency or keep a vessel open at or near the deploymentdiameter. The patency provided by the stent allows the stented segmentof the vessel to undergo positive remodeling at the increased deployeddiameter. By maintaining the patency of the stented segment at thisstage, the stent prevents negative remodeling. Remodeling refersgenerally to structural changes in the vessel wall that enhance itsload-bearing ability so that the vessel wall in the stented section canmaintain an increased diameter in the absence of the stent support. Aperiod of patency is required in order to obtain permanent positiveremodeling.

During this time period, the stent inhibits or prevents the naturalpulsatile function of the vessel. The stent structure prevents recoiland maintains a circular lumen while the vessel remodels and moldsitself to the stented diameter, which corresponds to positiveremodeling. Early recoil before sufficient modeling takes place canresult in negative remodeling, referring to molding of the stent to adiameter significantly less than the original stented diameter, forexample, 50% or less than the original deployment diameter.

As the polymer of the stent degrades, the radial strength of the stentdecreases and the load of the vessel is gradually transferred from thestent to the remodeled vessel wall. Remodeling of the vessel wallcontinues after loss of radial strength of the stent. Before the stentloses mechanical integrity, it is desirable for the stent structuralelements to become incorporated in the vessel wall by an endotheliallayer. The stent then breaks apart which allows vasomotion. The vesselwall continues to remodel as the vessel moves due to vasomotion. Thestent eventually erodes away completely leaving a healed vessel with anincreased diameter and which can exhibit vasomotion the same or similarto a healthy vessel section.

Poly(L-lactide) (PLLA) is attractive as a stent material due to itsrelatively high strength and a rigidity at human body temperature, about37° C. Since it has glass transition temperature between about 60 and65° C. (Medical Plastics and Biomaterials Magazine, March 1998), itremains stiff and rigid at human body temperature. This propertyfacilitates the ability of a stent to maintain a lumen at or near adeployed diameter without significant recoil.

PLLA has an in vitro degradation time of up to 3 years (Medical Plasticsand Biomaterials Magazine, March 1998; Medical Device Manufacturing &Technology 2005). The degradation time is the time required for completeloss of mass of a polymer construct, such as a stent. The degradationtime in vivo is shorter and depends on the animal model. In addition toan erosion profile, a PLLA stent has associated molecular weight andmechanical property (e.g., strength) profiles. As demonstrated above,the time dependence of the mechanical properties and mechanicalintegrity are important in the treatment of a diseased vessel. Thedegradation properties of PLLA do not generally coincide with what isrequired or desired for a particular treatment application. For example,it is desirable for a PLLA stent to have a degradation time of about twoyears (e.g., 22 to 26 months) for coronary vascular application, ofabout eighteen months (e.g., 16-20 months) for a peripheral application(e.g., superficial femoral artery (SFA)), and less than a year for nasalapplications. For coronary and peripheral applications, the radialstrength should be maintained for between one to six months, whereinmaintaining radial strength refers to the ability of the stent tomaintain the stented section at a diameter at least 50% of the originaldeployed diameter. For coronary and peripheral applications, the startof the loss of mechanical integrity should not occur until at leastbetween 2 and 4 months, where the start of the loss of mechanicalintegrity corresponds to breaking apart of structural elements.

It is desirable to adapt a PLLA stent to various applications, so thatit has the degradation behavior appropriate to each application, withoutsignificantly changing the composition of the stent. The embodiments ofthe present invention relate to adjusting the time-dependent degradationbehavior of a PLLA stent through inclusion L-lactide (LLA) monomer in astent body composed of PLLA.

The embodiments of the stent can include a body or scaffolding that iscomposed of PLLA with a targeted small amount of LLA monomer. The bodycan be composed of targeted amount of LLA and the rest of the body canbe 100% PLLA. Additionally, the body can be made of the targeted amountof LLA, additional components, and the rest of the body 100% PLLA. Thebody can be 95 wt % or greater and the rest of the body can be thetargeted amount of LLA and optionally the additional components. Theadditional components can be drugs, polymers, or filler materials suchas bioceramic particles. The present invention includes embodiments thatexclude additional components or exclude one or more of the additionalcomponents mentioned above.

In additional embodiments, the PLLA stent body can include additionallyor alternatively, d-lactide, meso-lactide, glycolide, lactic acid, ortheir oligomers with an Mn less than 1,000 g/mol. The concentration ofthese additional monomers can be the same as those disclosed forL-lactide.

The stent can further include a coating over the body or scaffolding. Inone embodiment, the coating can be a polymer and drug mixture. Forexample, the coating can be poly(D,L-lactide) and the drug could be anantiproliferative such as everolimus. The coating can be free of the LLAmonomer other than incidental migration or diffusion of LLA into thecoating.

Low molecular weight PLLA oligomers can also increase the degradationrate, and thus adjust degradation behavior. However, the increase isprimarily due to acidic end groups that act as catalysts to increasedegradation rate of the PLLA. Thus, the larger the oligomer, a higherthe weight fraction of oligomer in the PLLA that is required. Therefore,a much lower weight fraction of LLA monomer than given oligomer isrequired for a similar effect as the oligomer. The high weight fractionof oligomer can negatively impact the mechanical properties of thestent.

The qualitative and precise effect of LLA monomers on PLLA degradationbehavior is not known. For example, the amount of monomer required toobtain desired degradation behavior is not. This can be at leastpartially attributed to the fact that the degradation behavior of astent made from a semicrystalline degradable polyester, such as PLLA, isa complex function of several properties of the material and stent body.These properties include the intrinsic hydrolysis rate of the polymer(i.e., the chain scission reactions of the polymer backbone), the degreeof crystallinity, the morphology (size and distribution of crystallitedomains in the amorphous matrix), molecular weight (as measured by theinherent viscosity, number or weight average molecular weight), andstent body parameters (pattern, strut dimensions and the surface tovolume ratio).

A semicrystalline polymer generally may have insufficient strength andfracture toughness to provide adequate and safe treatment of a bloodvessel. The fabrication of the stent of present invention includesprocessing which increases the strength and fracture toughness of thefinal stent product. This processing provides certain characteristicsthat affect the degradation behavior, e.g., crystallinity, morphology.The strength and fracture toughness are increased by induced biaxialorientation of polymers in the hoop or circumferential and/or axialdirection, a particular range of the degree of crystallinity, and smalldispersed crystallites.

The stent is made from an extruded PLLA polymer tube that has beenradially expanded and axially stretched to provide the inducedorientation. The polymer tube is expanded by blow molding with a percentradial expansion between 200% and 500%, and a percent axial stretch from20% to 200%. The extruded PLLA tubing has a percent of axial stretchfrom 100% to 400%. The stent is formed from the expanded tube by lasercutting the tubing in its expanded state.

Additionally, the blow molding process is performed in a manner thatresults in small crystallites dispersed through an amorphous matrix.Prior to expansion, the tube is heated to a temperature between 70° C.and 95° C. to induce formation of smaller crystallites that enhancefracture toughness. The tube is quenched below the glass transitiontemperature (Tg) after expansion to prevent further crystal growth. Thedegree of crystallinity is 30-50%. Below 30% crystallinity the stentbody may not be have sufficient strength, while above 50% crystallinitythe stent body may be too brittle. The number average molecular weight(Mn) of scaffolding material in the final product (in g/mol) is between60,000 and 300,000, or more narrowly between 80,000 and 200,000.

An exemplary strut cross-section can be rectangular, for example,140×140 μm to 160×160 μm or having a cross-sectional area between 20,000and 25,000 μm².

In one embodiment of the present invention, the stent is composed ofPLLA and less than 1 wt % of LLA monomer. In a more preferred embodimentof the present invention the stent includes less than 0.9 wt %, lessthan 0.7 wt %, less than 0.5 wt %, less than 0.4 wt %, less than 0.3 wt%, less than 0.2 wt %, or less than 0.1 wt % of LLA monomer. In otherembodiments, the stent has between 0 and 1 wt % LLA, 1-2 wt % LLA, 2-3wt % LLA, 3-4 wt % LLA. However, as shown below, it is expected that anLLA content of greater than 1 wt % or greater than 2 wt % will result ina stent that does not maintain radial strength and mechanical strengthlong enough for effective treatment of a diseased section of a vessel.

The LLA monomer can be dispersed in the form of a powder or particularparticles through all or a portion of a stent body. The size of suchparticles can be less than 100 nm, between 100 nm and 200 nm, or greaterthan 200 nm, where size can refer to diameter or some othercharacteristic length. Alternatively, the LLA monomer can be mixed ordissolved on a molecular level with the PLLA.

It has been observed from in vitro and in vivo degradation studies ofPLLA stents with L-lactide monomer, discussed below, that LLA provides adramatic and unexpected increase in the degradation rate of the stent,particularly above about 1 wt %. Stents having monomer compositionsabove about 1 wt % LLA lose mechanical strength, lose mechanicalintegrity, and erode away in a fast way. Additionally, the lowconcentration of LLA are advantageous since the effect of the dispersedmonomer in the polymer has no or a minimal effect on the mechanicalproperties of the polymer.

Additionally, it is important that the LLA monomer be uniformly orsubstantially uniformly dispersed through the PLLA of the stent toprovide uniform degradation behavior through the body of the stent. Ithas been observed that for such low concentrations of LLA monomer, theuniformity of distribution is highly dependent on the manner of mixingor dispersing of the LLA. Thus, additional embodiments include a methodof mixing the LLA in the PLLA stent material.

In vitro and in vivo studies of degradation behavior can be used toassess the influence of LLA monomer concentration on the degradationbehavior of a PLLA stent. In addition, the influence can be assessedusing theoretical models.

A hydrolytic degradation model for aliphatic polyesters having the formMn(t)=Mn(0)exp(-Kt), wherein Mn(t) is the number average molecularweight at time t, Mn(0) is the number average molecular weight at t=0,and K is the hydrolytic degradation rate constant. Pitt, C. G., J. ofApplied Polymer Science 26, 3779-3787 (1981); Pitt, C. G., Biomaterials2, 215-220 (1981); Weir, N. A., Proceedings of the Institution ofMechanical Engineers, Part H: J. of Engineering in Medicine 218, 307-319(2004); Weir, N. A., Part H: J. of Engineering in Medicine 218, 321-330(2004). The assumptions inherent in the model are reasonable providedthat the mass loss has not occurred, since mass loss would affect theconcentrations of water and carboxylic end groups in the sample. Theequation can also be written as: ln[Mn(t)/Mn(0)]=-Kt. Therefore, byrepresenting data for Mn(t)/Mn(0) versus t on a log-linear plot, one mayinfer the hydrolytic degradation rate from the slope of the connectingpoints.

In vitro and in vivo degradation data has been generated for PLLA stentswith different concentrations of LLA monomer to investigate its effecton the degradation behavior of a PLLA stent. The in vivo data wasobtained using an animal model. Parameters and behavior that have beenmonitored include Mn, radial strength, appearance of cracks or fracturesin stent struts (structural/mechanical integrity), and degradation time.In all studies, the stent is processed as described herein. Theparameters of the stents are shown in Table 1 below. The stents includeda coating over a scaffolding composed of poly(DL-lactide) andEverolimus.

TABLE 1 Summary stent parameters in degradation studies SpecificationValue Backbone polymer (PLLA) Mw 180,000-200,000 Mn  90,000-100,000 Massof stent (18 mm length) 9.0 mg Mass/unit length 0.5 mg/mm Crystallinity 45% (as measured by DSC) Strut cross section 150 micron × 150 micronCoating thickness  3 microns Coating mass 308 μg (1:1polymer:Everolimus) Coating polymer Mw 66,000 Mn 39,000 Blow molding:Percent radial expansion 400% Percent axial elongation  20% Lasermachining 120 fs laser

The in vitro studies were performed in a phosphate buffered saline (PBSbuffer) solution at 37° C. The Mn of the stent was measured by GPC usingpolystyrene standards. The animal model for the in vivo studies wasYucatan mini swine.

In the in vitro study, the Mn dependence on time is plotted in the formof ln[Mn(t)/Mn(0)] versus time to assess the predictive ability of thehydrolytic degradation model. The model is then used to assess theeffect of LLA on the degradation of PLLA.

In the in vitro studies, the Mn as a function of time was also measuredfor PLLA stents having different concentrations of LLA monomer. In thesestudies, the LLA monomer was added to the PLLA resin in a mannerdescribed herein. FIG. 2 depicts a plot of ln[Mn(t)/Mn(0)] versus timefor this set for four groups of stents: nominal (0.05 wt % LLA), 0.2 wt% LLA, 0.55 wt % LLA, 1 wt % LLA, and 3.3 wt % LLA monomer in PLLA.

The data for each concentration is fitted with a straight line, theslope of which provides the rate constant, K, of the degradation model.The rate constant, K, for each concentration of LLA monomer issummarized in Table 2 and FIG. 3.

TABLE 2 Rate constant K for degradation model from in vitro data. StentGroup: wt % LLA k (×10³) (days⁻¹) R2 Nominal 1.9 0.962 0.2 3.0 0.9720.55 7.2 0.969 1 13.4 0.960 3.3 48.8 0.989The coefficient of determination, R2, is given for each group of data toassess the predictability of the linear degradation model. The closer R2is to 1, the more reliable the predictive ability of the model. The datain Table 2 shows an unexpectedly dramatic effect of LLA monomer contenton the degradation of the stent. For example, when the LLA concentrationincreases from near 0 to 0.55 wt %, K increases near three times andfrom near 0 to 1 wt %, the rate constant increases near six times. Thedifferences in the change in the molecular weight with time as the LLAconcentration increases is just as dramatic.

Table 3 shows the percent drop in Mn and the Mn at time points based onthe model predictions for 1 wt % LLA for the second set of data.

TABLE 3 Change in molecular weight with time for PLLA with 1 wt % PLLADays Drop in Mn Mn 20 23% 77k 40 41% 59k 60 55% 45k 80 65% 35k 120 80%20kAs indicated above, the loss of mechanical integrity starts before massloss. In addition, the loss of radial strength occurs before or alongwith loss of mechanical integrity. At the entanglement molecular weight,about 17,000, the polymer no longer has mechanical properties anddisintegrates under a applied load. The stent loses mechanical integritywell before the entanglement molecular weight.

As indicated above, the in vitro start of mass loss is expected to be anupper limit for in vivo degradation, that is, mass loss starts at anearlier time in vivo. In addition, the start of mechanical integrityloss and the loss of radial strength might also occur at an earlier timein vivo than in vitro. Table 4 shows in vivo and in vitro data for massloss for a PLLA stent with no LLA monomer. The difference in the timethat mass loss starts and the difference in the degree of loss issignificant.

TABLE 4 Comparison of in vitro and in vivo mass loss for PLLA stent withno LLA monomer. Time In vitro In vivo (months) mass loss Mass loss 9 0%20% 12 0% 30% 18 — 35% 21 9% —

For coronary and peripheral applications, it is believed that themechanical integrity should remain intact for at least 2 to 4 monthswithout severe fractures (e.g, breaking of struts) after implantation toallow incorporation of stent into vessel wall. Additionally, it isbelieved that radial strength should be maintained for at least about 1month to prevent negative remodeling. The radial strength is expected tobe lost prior to the mechanical integrity and the start of the loss ofmechanical integrity is expected to start before mass loss. A prelude tothe start of the loss of mechanical integrity corresponds to theformation of cracks in the stent. Therefore, based on this in vitrodata, the LLA concentration should be less than 1 wt % for the stent tomaintain radial strength and mechanical integrity for the desired timesfor coronary and peripheral applications.

In another set of in vitro studies, stent groups with near 0 wt %, 0.2wt %, 0.55 wt %, 1 wt % and about 3 wt % of LLA in PLLA were studied andtracked over a four month period. In these studies, the radial strengthand stent integrity were tracked. Table 5 summarizes the observed changein radial strength and mechanical integrity after radial strengthtesting for the different groups. As shown in Table 5, for near 0 wt %,0.2 wt %, and 0.55 wt % LLA, the radial strength is maintained up tomore than 4 months and there were no fractures observed up to more than4 months. For 1 wt % LLA, the radial strength showed a consistentdecrease between about 1½ months to about 3 months. The drop in radialstrength might occur sooner in vivo. Additionally, the significantfractures at only 42 days indicate an early loss of radial strength andmechanical integrity. These results indicate a stent with greater than 1wt % LLA might not be a great candidate for coronary or peripheralapplications. The drop in radial strength and significant fractures inthe stent with about 3 wt % LLA indicate that a PLLA stent with thisconcentration is totally unsuitable.

TABLE 5 Radial strength and mechanical integrity of PLLA stents withdifferent LLA monomer content. LLA content (wt %) Radial strengthMechanical integrity 0.05 maintained up to more no fractures at up tomore than than 4 months (126 4 months (126 days) days) 0.2 maintained upto more no fractures at up to more than than 4 months (126 4 months (126days) days) 0.55 maintained up to more no fractures at up to more thanthan 4 months (126 4 months (126 days) days) 1.0 consistent decreasestent segments in pieces observed between ~1½ months at 42 days to about3 months ~3 consistent decrease stent segments in pieces observedstarting at about 2 at 14 days weeks

Preclinical in vivo (animal) results for stent groups with a 0 wt %, 0.1wt %, 0.4 wt % LLA, ca. 0.6 wt % LLA, 1 wt % LLA, and 3.8 wt % LLA havebeen obtained for up to 28 days after implantation. For stents with 0.4wt % LLA and ca. 0.6% there were no fractures observed at 28 days afterimplantation. For stents with 1 wt % LLA, fractures were observed at 28days after implantation. For stents with 3.8 wt % LLA, there weresignificant fractures at only 7 days and the stent broke into piecesafter 28 days.

It has been observed by the inventor that forming a PLLA tube intendedto have between 0.05-0.5 wt % LLA by mechanical blending of therequisite amount of LLA in PLLA fails to result in tubes with uniformblending of the LLA. Large variations in LLA concentration betweenstents formed from tubes in this manner were observed.

In some embodiments, the LLA can be mixed with PLLA using solventblending. Solvent blending may include making a master batch of PLLA andLLA mixture with concentrations of LLA larger than the targetconcentration. The master batch is made by dissolving LLA and PLLA in asolvent such as chloroform. The chloroform is evaporated to form themaster batch which is a uniform mixture of PLLA and LLA. The masterbatch is then mixed by melt processing, such as in an extruder, with asufficient amount of PLLA to obtain the target LLA concentration. Themethod is illustrated by the following example:

-   Step 1: Dissolve 2 g LLA monomer and 8 g PLLA in 400 ml chloroform-   Step 2: Evaporate chloroform to form uniform mixture of PLLA and LLA    with 25 wt % LLA.-   Step 3: Blend 25 wt % LLA mixture in extruder with 390 g of PLLA to    obtain 0.5 wt % LLA in PLLA.

In other embodiments, LLA can be mixed with PLLA using spray coating.The spray coating can include preparing a uniform mixture of LLA in PLLAincludes dissolving LLA in a solvent such as methanol to form a solutionand spraying the solution on PLLA pellets. The solvent is removedleaving the LLA deposited on the PLLA pellets. The PLLA pellets are meltprocessed, in an extruder, for example, to form tubes with the targetconcentration of LLA in PLLA. The method is illustrated by the followingexample:

-   Step 1: Dissolve 5 g LLA in 100 ml anhydrous methanol to form    solution.-   Step 2: Spray solution onto 1 kg of PLLA pellets and stir.-   Step 3: Place pellets in vacuum oven to remove solvent.-   Step 4: Place pellets in extruder and form tube with 0.5 wt % of    LLA.

At low concentrations of LLA (e.g., less than 1 wt %), melt blending canresult in nonuniform mixing of LLA monomer and PLLA polymer.Specifically, phase separation of LLA from PLLA occurs and has beenobserved by the inventors. The phase separation is likely due to thelarge density variation between LLA monomer and PLLA. The nonuniformdistribution may result in variations in the degradation rate of thefinal products made from the PLLA.

To address this, further embodiments of making a PLLA material withuniform LLA monomer content for making a stent or other medical deviceinvolves control of polymerization conditions during PLLA synthesis,control of post-processing conditions, or both. Such embodiments canprovide PLLA pellets for processing into a stent with uniformlydistributed LLA monomer at both low and high concentrations of LLA,i.e., from 0.01 wt % to 5 wt %.

In conventional medical grade PLLA polymer material synthesis, it is abasic goal or requirement to obtain high purity PLLA resin with very lowLLA monomer content. Polymerization is performed with the goal tominimize the LLA monomer content in the final product since it isgenerally believed that large amount of monomers may cause irritation invivo for some applications. Also, it is generally common sense that anyimpurities may complicate the downstream processes depending on theapplications. The polymerization temperature and reaction time are twovariables of the polymerization process that significantly influence theconversion of LLA to PLLA, and thus the LLA content in a PLLA product.

An example of PLLA synthesis is ring-opening polymerization of L-lactidein the bulk phase using stannous octoate(tin(II1 2-ethylhexanoate) ascatalyst. (S. Hyon et al., Biomaterials 16 (1997) 1503-1508) Hyon et al.have presented the monomer conversion of PLLA as a function of time forpolymerizations carried out over a temperature range 120° C. to 220° C.FIG. 4_(FIG. 4 of Hyon et al.) depicts the dependence of the monomerconversion on the reaction time for bulk polymerization of L-lactide atdifferent temperatures (catalyst=0.05 wt %). FIG. shows the LLA monomerconversion to the resulting PLLA increases with time in an initial stageof polymerization at all temperatures. In this initial stage, theconversion of the resultant polymers exhibits almost a linear increaseup to the monomer conversion of about 80%. In later stages, or longerreaction times, there is a gradual decrease in the monomer conversionwhich becomes more pronounced at higher temperatures, i.e., above 180°C.

In order to meet the requirement of low LLA monomer content,polymerization is conventionally performed at relatively lowtemperatures (lower than 180° C., preferably 140-160° C.) with shortreaction times (less than 16 h). The lower temperatures and shortreaction times reduce or prevent depolymerization that regeneratesmonomers. Sometimes the reaction temperature at a later polymerizationstage is lowered to eliminate any depolymerization.

In general, polymerization temperatures for PLLA are between 120-180°C., preferably, 140-160° C. The molecular weight of the product is alsoaffected by the temperature and reaction time, but is also controlled bythe monomer/initiator ratio. Typical initiators are alcohols likedodecanol or hexanol. To obtain a PLLA product with weight averagemolecular weight between 200 to 600 kg/mole, a monomer/initiator molarratio of about 1500:1 to 4000:1 (molar) may be used. Additionally, amonomer/catalyst weight ratio of between 1000:1 to 2000:1 (weight) isused. After polymerization, the final product which is cut into quitesmall pellets, for example, a pellet can be about 0.1 g.

Additionally, residual monomer content is reduced by subjecting thepolymer to post-processing that extracts monomer from the synthesizedpolymer. There are several monomer extraction post-processing methodsthat are used such as solvent extraction, supercritical CO₂ extraction,and vacuum/high temperature extraction. In solvent extraction, thepolymer is soaked in a solvent for the monomer which is a nonsolvent forpolymer. In CO₂ extraction, the PLLA material is exposed tosupercritical carbon dioxide which extracts LLA monomer from thematerial. In vacuum/high temperature extraction, a PLLA sample isexposed to temperatures in a range between 140-160° C. at very lowpressures, less than 0.1 atm. There are several process variables ineach of these processes that influence the degree of monomer extraction.However, the primary variable is time of exposure to the process. Thelonger the exposure, the higher the removal of monomer.

Extraction solvents for extracting LLA from PLLA include methanol,ethanol etc, which is solvent for LLA but poor or non-solvent for PLLA.The amount extracted is a function of soaking time in the solvent.Different soaking times result in final products with different monomercontent. The higher the soaking time, the higher the removal of themonomer and the lower the monomer content of the resulting product. Dueto the small size of the pellets, the extraction results in uniformcontent of monomer in the pallet. As an example, for a starting LLAmonomer concentration of 5 wt %, the exposure time necessary to get atarget monomer concentration less than 0.1% in solvents such as methanolmay be at least 24 h.

Thus, conventionally, polymerization is performed to result in a productwith a monomer content as low as possible. Additionally, monomerextraction conventionally is applied to remove all monomer or as much aspractically possible from a polymer.

In the various aspects disclosed, polymerization, alone or incombination with post-processing extraction, is performed todeliberately result in a PLLA polymer material with a LLA monomercontent higher than conventionally desirable with the object to provideor control the degradation rate of a stent made from the PLLA material.Additionally, post-processing monomer extraction may be applied eitheralone or in combination with polymerization to provide a PLLA productwith such monomer content. In some embodiments, the target LLA monomerconcentration resulting from polymerization, post-processing, or both isbetween 0.01 to 2 wt %, 0.05 to 5 wt %, 0.05 to 2 wt %, 0.05 to 1 or wt% or 0.5 to 2 wt %.

In other embodiments, the target LLA monomer concentration is less than2 wt %, less than 1 wt %, less than 0.9 wt %, less than 0.7 wt %, lessthan 0.5 wt %, less than 0.4 wt %, less than 0.3 wt %, less than 0.2 wt%, or less than 0.1 wt % of LLA monomer. In additional embodiments, thetarget LLA monomer concentration is 0 and 1 wt % LLA, 1-2 wt % LLA, 2-3wt % LLA.

In certain embodiments of the invention, the target content of LLA ofthe PLLA product is controlled through control of the post-processing.In particular, the LLA monomer content is controlled through theexposure time of the PLLA pellets in a process such as solventextraction, supercritical CO₂ extraction, or vacuum/high temperatureextraction. For solvent extraction, the exposure time corresponds to thetime the PLLA pellets are soaked in the extraction solvent. Forsupercritical CO₂ extraction, the exposure time corresponds to the timeof exposure of the PLLA pellets to supercritical CO₂ at the extractionconditions. For vacuum/high temperature extraction, the exposure timecorresponds to the time of exposure of the PLLA pellets to thevacuum/high temperature extraction conditions.

In these embodiments, a PLLA polymer material is obtained that has arelatively high LLA monomer content. The LLA content may be, forexample, greater than 3 wt %, greater than 4 wt %, or greater than 5 wt% LLA. The PLLA is exposed to the extraction conditions for a targetexposure time during which LLA is extracted from the PLLA and thenextraction is terminated. Terminating extraction can correspond toremoving the PLLA from the solvent or extraction chamber or changing theconditions of an extraction chamber to conditions that do not extractLLA. In other embodiments, the extraction can be performed in stages inwhich extraction can be terminated before reaching a target LLA contentand then exposed one or more additional times until a target LLA contentis reached.

After extraction is completed, the PLLA has a target LLA content due tothe extraction of the LLA from the exposure. The target exposure time isthe exposure time that results in a target LLA monomer content, rangesof which are provided above, in the PLLA polymer material. The targetexposure time can be determined from exposure of a plurality of PLLAsamples at different exposure times. For example, an initial exposuretime for a sample can be selected and then additional samples can beexposed at progressively higher exposure times until an exposure isreached that provides the target LLA monomer content.

In additional embodiments, a PLLA material with the target LLA contentis obtained through control of conditions of post-processing monomerextraction in combination with control of polymerization. In suchembodiments, a PLLA polymer material with relatively high monomercontent is prepared first by control of polymerization conditions. PLLApolymer material refers to PLLA polymer, LLA monomer, and any otherimpurities or residue.

Making a PLLA polymer material from polymerization with such LLA monomercontent can be accomplished with various combinations of reaction timesand temperatures. The PLLA polymer material can be made with a LLAmonomer content of, for example, greater than 3 wt %, greater than 4 wt%, or greater than 5 wt % LLA. In one embodiment, the relatively highLLA monomer content can be achieved with a relatively low reactiontemperature and short reaction time. In another embodiment, therelatively high LLA monomer content can be achieved with high reactiontemperature and relatively long reaction time. In the third embodiment,the relatively high LLA monomer content can be achieved in two steps.During a first stage where a high consumption of LLA is achieved with arelatively low reaction temperature and long enough reaction time, andthen increasing reaction temperature at a late polymerization stage andprolonging reaction time sufficiently to obtain a desired LLA monomercontent. For example, the reaction temperature can be less than 160° C.for a first stage with a reaction time of less than 20 h. The LLAmonomer content at the end of this first stage can be lower than 1%,lower than 0.5%, or lower than 0.1%. The first stage is then followed bya second stage with a temperature greater than 180° C. or 180-200° C.and a reaction time of more than 4 h. The LLA monomer content of thePLLA polymer material at the end of the second stage may be greater than4 wt %, greater than 5 wt %, 5-7 wt %, 7-10 wt %, or greater than 10 wt%.

The PLLA polymer material at the end of the second stage with therelatively high monomer content can then be subjected to post-processingextraction to reduce the LLA content to a target LLA monomer content.The LLA monomer content can be reduced by the extraction to any value ofthe disclosed target LLA content ranges, or any value between 100 ppm to5 wt %.

In further embodiments, a PLLA polymer material with the target LLAcontent can be made by a polymerization scheme that includes two stages.In the first stage, the polymerization has a relatively high conversionresulting in a PLLA material with a relatively low LLA monomer content.The relatively low LLA monomer content can be lower than the disclosedtarget LLA monomer content ranges or within such ranges. In the secondstage, the reaction temperature is increased to cause depolymerizationwhich increases the LLA monomer content to a target LLA monomer content.

For example, the conversion of the first stage polymerization can begreater than 98%, 99%, or 99.5%. The LLA monomer content at the end ofthe first stage may be less than 2%, 1%, or 0.5%. In the second stage,the depolymerization causes the LLA monomer content to increase to anyrange or any value in the disclosed ranges of the target LLA monomercontent.

The resulting PLLA material from the depolymerization stage may then beused for making a stent without any post-processing extraction processto decrease the LLA monomer content. Optionally, post-processingextraction can be performed on the PLLA material to reduce the LLAmonomer content further to a lower target LLA monomer content.

In the first stage, the polymerization that achieves the high conversioncan be achieved using a variety of combinations of temperature andreaction times. The polymerization can be performed at a temperature ortemperatures between 120-160° C., 125-150° C., or 140-150° C. Thereaction time is sufficient to provide the high conversion that isdesired, for example, the reaction time may be greater than 10 h, 15 h,20 h, or more than 20 h. These reaction time ranges may correspond toany of the above temperature ranges. The reaction temperature for thesecond stage may be greater than 180° C., 180-200° C., or greater than200° C. The reaction time is sufficient to increase the LLA monomercontent to a target LLA monomer content. For example, the reaction timemay be less than 6 h, less than 1 h, 1-2 h, 2-3 h, 3-4 h, or 4-6 h.

In additional embodiments, a PLLA material with a target LLA monomercontent can be made by a polymerization reaction performed at a selectedreaction temperature for a selected reaction time to results in the PLLAmaterial with the target LLA content. In this embodiment, thepolymerization reaction provides the PLLA material without furtherpost-processing extraction of LLA monomer. Optionally, post-processingextraction can be performed on the PLLA material to reduce the LLAmonomer content further to a lower target LLA monomer content.Embodiments include any temperature between 120-180° C. and any reactiontime between 1 h and 100 h and any combination of these temperatures andreaction times that provides a PLLA polymer material with any of theranges and values with in such ranges of the disclosed target LLAmonomer content.

The selected reaction temperature and reaction time that provides atarget LLA monomer content may be determined in at least two ways. Oneway is to fix a reaction temperature and perform the polymerizationreaction at different reaction times to determine the reaction time thatprovides the target LLA monomer content at the selected reactiontemperature. For example, a target LLA monomer content may be 0.5% wt %and a selected reaction temperature may be 140° C. The polymerizationreaction may then be performed for reaction times between at least 6 hand 30 h. The selected reaction time is that which corresponds to thereaction time that provides the target LLA monomer content of 0.5 wt %.

Another way to determine the selected reaction temperature and reactiontime that provides a target LLA monomer content is to fix a reactiontime and perform the polymerization reaction at different reactiontemperatures to determine the reaction temperature that provides thetarget LLA monomer content in the selected reaction time. For example, atarget LLA monomer content may be 0.5 wt % and a selected reaction timemay be around 20 h. The polymerization reaction may then be performedfor reaction temperatures between at least 120 and 170° C. The selectedreaction temperature is that which corresponds to the reactiontemperature that provides the target LLA monomer content of 0.5 wt %. Ifthe reaction time is too short, there may not provide the target LLAmonomer content (see FIG. 4). Additionally, there is a lower limit ofthe LLA content at higher temperatures due to depolymerization (see FIG.4).

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1. (canceled)
 2. A method of making a stent for treating a diseasedsection of a blood vessel, comprising: extruding a polymer compositioncomprising poly(L-lactide) to form a tube; processing the extruded tubeto increase crystallinity of the tube to between 20-50%; and lasercutting the processed tube to form a stent comprising a scaffolding,wherein the scaffolding comprises greater than 95 wt % poly(L-lactide)and between 0.05 to 0.5 wt % L-lactide monomer mixed, dispersed, ordissolved within the poly(L-lactide).
 3. The method of claim 2, whereinthe processing comprises heating and radially expanding the extrudedtube to induce orientation in the polymer.
 4. The method of claim 3,wherein extruded tube is radially expanded to a degree of radialexpansion of between 200-500%.
 5. The method of claim 3, wherein theextruded tube is heated to a temperature between 70° C. and 95° C. priorto expansion.
 6. The method of claim 2, wherein the scaffolding iscomposed of a pattern of struts, the pattern including a plurality ofcylindrical rings connected by linking struts.
 7. The method of claim 6,wherein the struts have a rectangular cross-section, the cross-sectionalarea being between 20,000 and 25,000 μm².
 8. The method of claim 2,wherein the number average molecular weight of the poly(L-lactide) ofthe scaffolding is between 60,000 and 300,000, relative to polystyrenestandards.
 9. The method of claim 2, further comprising mixing theL-lactide monomer with the poly(L-lactide).
 10. The method of claim 9,wherein the mixing comprises mechanically blending the L-lactide monomerwith the poly(L-lactide).
 11. The method of claim 9, wherein theL-lactide monomer is mixed with the poly(L-lactide) using solventblending.
 12. The method of claim 11, wherein the solvent blendingcomprises: making a master batch of the poly(L-lactide) and theL-lactide monomer having a concentration of the L-lactide monomer largerthan a target concentration, wherein the master batch is made bydissolving the L-lactide monomer and the poly(L-lactide) in a solventand evaporating the solvent to form the master batch which is a mixtureof the poly(L-lactide) and the L-lactide monomer; and making the polymercomposition by mixing the master batch with an amount of thepoly(L-lactide) that is sufficient to obtain the target L-lactidemonomer concentration in the polymer composition.
 13. The method ofclaim 2, further comprising: preparing a solution of the L-lactidemonomer dissolved in a solvent; spraying the solution on pellets of thepoly(L-lactide); removing the solvent to leave the L-lactide monomerdeposited on the pellets; and using the polymer with the depositedL-lactide monomer as the polymer composition for the extruding step.